Magnetic resonance (MR) imaging benefits from a static magnetic field that is stable over time. The main (or static) magnetic field of the MRI scanner is commonly denoted as the B0 magnetic field, and has a high value to align nuclear spins (in a statistical sense). In some MR scanners used for medical imaging, B0 is in the range 0.2 Tesla to 3.0 Tesla, and even higher values, e.g. B0=7 Tesla, are used in research applications. Superconducting magnets are generally used to achieve these high magnetic fields. Some suitable superconducting materials for fabricating the superconducting magnet windings include niobium-titanium, niobium-tin, or so forth which have a critical temperature (TC) that is typically below 20K. Thus, the superconducting magnet windings are immersed in liquid helium (LHe) contained in a vacuum-jacketed LHe dewar or are disposed in some other type of cryostat to maintain the windings at suitably low cryogenic temperature. In magnets employing superconducting materials with higher TC, the magnet cryostat can take other forms such as being immersed in liquid nitrogen (LN2) or vacuum.
However, small temporal B0 variations on the order of only a few nT can degrade the MR image quality. The amount of nT variation that can be tolerated in typical medical imaging applications depends on the frequency and ranges from about 1 to 100 nT at 0.01 to 100 Hz. The tolerable variation is thus in the parts-per-billion (ppb) range. B0 variation larger than this can easily be caused by external sources around the MR scanner, such as electricity lines or moving magnetic objects like trains, cars, elevators in the neighborhood of the scanner. Therefore, MR magnets typically are provisioned for compensating external field variations in order to have good image quality. Such a provision is referred to herein as a B0 compensation system. The design goal for a B0 compensation system is typically to reduce external B0 magnetic field disturbances by a factor 10 to 100. The B0 magnetic field disturbances reduction factor is referred to herein as the shield factor—a higher shield factor corresponds to better B0 compensation. (The “external” B0 field refers to the B0 field outside of the magnet itself, and typically refers to the B0 field in the imaging field-of-view (FOV), e.g. at the isocenter of the bore of a horizontal-bore type MR scanner).
Various types of B0 compensation systems have been developed. In active compensation approaches, a magnetic field sensor is installed in or near to the imaging FOV, and the measured magnetic field is used for feedback control to actively counter B0 magnetic field disturbances. One way to actively compensate for B0 variation is to actively apply a compensating magnetic field. For example, control electronics may drive one or more coils to generate a compensating field at the magnet. Alternatively, since the magnetic resonance frequency is proportional to the magnetic field (with the gyromagnetic ratio serving as the proportionality constant), the active compensation can be a frequency adjustment performed on the measured MR signals (e.g., in software).
Other types of B0 compensation systems are passive, and entail adding superconducting circuitry disposed with the superconducting magnet windings in the magnet cryostat. Some passive B0 compensation systems are described in, e.g. Reichert, U.S. Pat. No. 4,926,289 (“Actively shielded, superconducting magnet of an NMR tomography apparatus”) and Overweg, U.S. Pat. No. 5,426,366 (“Magnetic resonance apparatus comprising a superconducting magnet”). These designs are based on the principle that a superconducting circuit keeps its magnetic flux constant. In one design, the B0 compensation circuit is electrically connected with the magnet windings, with the connection made at strategically chosen locations, typically within winding coils. By designing the connection point properly, it is possible to obtain a shield factor of more than 100. In an alternative design, magnetic coupling of windings of the B0 compensation circuit with the magnet windings is substituted for the electrical connection. Again, with proper coupling design a shield factor of more than 100 can be obtained for quasi-DC disturbances.
It is recognized herein that these existing B0 compensation systems have certain disadvantages. In the case of active B0 compensation, sufficiently accurate B0 measurements (accuracy in the ppb range) are difficult to obtain due to factors such as difficulty in positioning the magnetic field sensor close to (and preferably symmetric respective to) the imaging FOV. In the case of passive B0 compensation, the gain is fixed and cannot be adjusted for less homogenous disturbance sources. It is also not generally feasible to compensate for disturbances that are not-quasi DC. This type of disturbance has a fixed frequency response. The metallic cryostat of the magnet influences the response of the passive B0 compensation system and this cannot be corrected because the compensation is passive and non-adjustable. In the case of a passive B0 compensation system with electrical connection to the MR magnet, the optimal connection points are usually inside of coil windings of the MR magnet, requiring extra lead-in-lead-out connections at the coil which adds manufacturing cost and complexity. In the case of a passive B0 compensation system with inductive coupling to the MR magnet, the compensation circuit requires extra wound superconducting coils, which must be designed to meet the shield factor design basis (e.g. a shield factor of at least 100 in some designs), which again increases cost and manufacturing complexity.
The following discloses a new and improved systems and methods that address the above referenced issues, and others.